Histology is currently the most accurate way to study tissue morphology even though the shape and the size of tissue components do not necessarily remain intact during the fixation, embedding, and sectioning processes involved in histology. Histology has the disadvantage that it requires tissue samples to be removed from patients in biopsy procedures.
Optical methods for studying tissue morphology in vivo have the advantage over histology that biopsies are not required. Optical coherence tomography (OCT) is an interferometric technique for obtaining images depicting sub-surface tissue morphology. OCT images can have axial resolution of less than 10 μm and may image tissues to depths of more than 1 mm. OCT can be used to study high-risk tissue sites without performing unnecessary biopsies and tissue removal [8]. Micro-invasive carcinoma can be distinguished from normal bronchial epithelium using epithelium thickness information measured by OCT [9]. However, compared to histology, current OCT techniques do not provide as detailed structural information about certain tissue components such as smooth muscle and different types of collagen.
A. Polarization-Sensitive OCT
Interrogating tissue by polarized light, polarization-sensitive OCT (PS-OCT) can provide additional information about birefringence properties of tissue. PS-OCT may provide better differentiation between selected tissue components as compared to polarization-insensitive OCT imaging.
Early PS-OCT systems were implemented by Michelson interferometers in free space using bulk optical components [4-11]. Controlling the polarization state of light and obtaining stable polarization states are much more feasible in free-space interferometers than in fiber-based interferometers since the polarization state of light does not change as it propagates in free space. In contrast, single mode fibers do not preserve the polarization of light due to the fiber birefringence associated with any deviation from fiber circular symmetry. Despite this disadvantage, fiber-based interferometers are much more tolerant to alignment and handling issues than free-space interferometers, offering the possibility of more robust systems for clinical use. In addition, the difficulties associated with implementing circulators with bulk optical components restrict the implementation of free space interferometers to cases where the signal to noise ratio (SNR) is high.
C. E. Saxer, et al., [12] used a fiber-based PS-OCT system to image burned tissue in vivo to determine the burn depth [13]. Saxter et al. calculated the retardation of a sample even though the exact incident polarization states on the sample were unknown.
M. C. Pierce, et al., [14] and B. H. Park, et al., [15, 16] showed that tissue birefringence and optics axis can be determined by data from alternating the polarization states of incident beam for successive A-lines at two polarization states perpendicular in a Poincaré sphere. However, this approach requires oversampling and restricted lateral scanning speed since the two-state polarization interrogation needs to be carried out at (nearly) the same location [17, 18]. Also, in an endoscopic PS-OCT system based on this method, sample arm motions affected the measurements [19].
Frequency multiplexing can be used to simultaneously measure the reflectance of two input polarization states, overcoming issues associated with sample arm motion or birefringence changes of the optical probe due to the probe bending or rotation [17, 20]. This method increases system cost and requires elaborate synchronization.
B. Baumann, et al., [18] reported an alternative approach to multiplexed PS-OCT using a passive polarization delay unit. This system was capable of operating at faster A-line scanning rates and did not require complex synchronization.
Polarization-maintaining (PM) fibers have been used to build PS-OCT systems [21-25]. However, due to the large birefringence of PM fibers, the lengths of the reference and sample arms' paths had to be exactly equal or additional numerical processing is required [23, 24].
Rotary and rotary-pullback fiber optic probes are widely used scanning mechanisms for generating 2-dimensional and 3-dimensional OCT images, respectively, of cylindrically symmetric structures. These probes are commonly made using single mode (SM) optical fiber and driven from the proximal end using rotary motors. Flexible torque-transmitting elements such as speedometer cables are used to transmit the rotary motion to the distal imaging tip. As the spinning SM fiber is continuously flexing and in motion, the polarization state of the light being emitted from the tip is constantly varying. Since the reference arm in a Mach-Zehnder interferometer does not share a common path with the sample arm the varying polarization state introduced by the rotating probe in combination with the fixed polarization state of light in the non-moving reference arm tends to create imaging artifacts.
One approach to mitigating the polarization effects of a rotating fiber optic sample arm is to use a common path probe where a partial reflection at the distal tip of the probe serves as the reference arm [36, 37]. However, being a Michelson type of OCT interferometer, this type of probe has less sensitivity relative to Mach-Zehnder interferometers [38]. An alternative approach to compensate for polarization effects in rotating probes is to use a polarization-diversity detection (PDD) scheme [39, 40]. In this scheme, polarization beam splitters are used to separate an interference signal into orthogonal polarization states. The reference beam power is equalized between the two polarization states and an image may be made from the square root of the sum of the squares of the intensities of the polarization states.
PDD schemes have been realized in the literature using free-space optics configurations [20, 41]. However, this type of setup is costly, difficult to miniaturize, and cumbersome since it involves the alignment of multiple beamsplitters and collimators. A fiber-based PDD OCT system is commercially available (PSOCT-1300, Thorlabs Inc.). However, balancing this system requires iterative adjustment of up to four polarization controllers and does not necessarily converge rapidly to an acceptable solution.
Autofluorescence (AF) bronchoscopy has proven to be more sensitive (up to 6 times) for detecting intraepithelial neoplastic lesions than white-light bronchoscopy [42-45]. AF imaging can provide valuable information about biochemical properties of tissue. AF imaging used in combination with OCT imaging (AF-OCT imaging) can provide biochemical information co-localized with structural information. The complementary nature of these two types of information makes AF-OCT imaging interesting for diagnostic applications. For instance, AF-OCT imaging can be used to study how disease processes change the structural as well as biochemical properties of airway tissue.
An AF-OCT system can be used to estimate AF loss due to epithelial scattering and absorption since the AF signal measured at the epithelial surface includes epithelial scattering and absorption effects [48]. Correcting AF intensity to remove the effects of absorption and scattering introduced by varying epithelial thickness determined by OCT can identify the contribution of submucosa fluorophores to the AF signal. OCT can measure epithelium thickness directly and determine which portion of the AF signal loss may be attributed to additional epithelial scattering.
OCT systems are typically designed to operate in the near infrared wavelength range where there is a good balance between tissue penetration and resolution. However, AF imaging systems are usually designed with visible or UV light sources to access biological chromophores. Optical components capable of operating in these two very different wavelength ranges are required to combine the two modalities. Also, endoscopic imaging of cylindrically symmetric structures such as airways typically uses rotary fiber optic probes. Combining AF and OCT is particularly challenging when imaging endoscopically via fiber optic probes. Several articles have been published on methods for combining OCT and AF imaging.
A combined AF-OCT imaging system can be obtained by combining two separate imaging systems either in free space using bulk optical components and dichroic mirrors [49-54], or by using a probe with separate adjacent fibers for the two modalities [55-61]. The former approach is unsuitable for imaging hard-to-reach places such as airways. The latter approach compromises the co-registration of the two modalities.
Other reported approaches used single Ti-Sapphire broadband femtosecond (fs) laser sources for both spectral domain OCT and multiphoton excited fluorescence imaging systems [63-70]. Using double-clad fibers (DCF) can provide a common path for the two modalities, ensuring co-registration [71-78]. Also, owing to the large inner cladding, DCFs have proven to be effective for the collection of AF emission photons [76,77].
There remains a need for robust practical and cost-effective systems for performing OCT and/or fluorescence-based imaging. There is a particular need for such systems capable of imaging narrow airways in the lung. A practical and cost-effective imaging system that combines OCT imaging (optionally PS-OCT) and fluorescence imaging (optionally AF imaging) would be highly beneficial for use in cancer screening and treatment as well as other medical applications. Such a system could be applied to guide biopsies and/or to make biopsies unnecessary in some circumstances.